A magnetic resonance imaging (MRI) system generates a static magnetic field B0. The magnetic field B0 produced by the MRI system will suffer from a degree of inhomogeneity. The homogeneity of the magnetic field B0 may vary due to design limitations of the MRI system, or compromises made in the manufacturing of the MRI system. A patient being imaged by an MRI system introduces further inhomogeneity to the B0 field due to variations in tissue composition within a region of tissue being imaged. For example, the degree of inhomogeneity may increase due to susceptibility differences at tissue boundaries within the region of tissue being imaged. The degree of inhomogeneity may also increase due to physiological processes that may occur within the region of tissue being imaged during an MRI procedure, including respiration, cardiac motion, and other physiological processes. These B0 field inhomogeneities may cause a variety image artifacts which reduce the image quality of the image acquired. As the B0 field strength increases, (e.g., from 1.5 T to 7 T) these issues may become more severe.
To address the inhomogeneities in the static B0 field, an MRI system is fitted with additional B0 field correction coils to correct the inhomogeneities in the B0 field. This is called “shimming” and the field correction coils are called B0 shimming coils. There are several approaches to implementing B0 shimming coils. A first conventional approach is to build B0 shimming coils into the MRI system itself. Most commercial MRI systems have B0 shimming coils built into the MRI system near the system gradient coils. This approach may be termed “global shimming”. A disadvantage to this global shimming approach is that the B0 shimming coils built into the MRI system near the system gradient coils are large and relatively far away from anatomy being imaged, and thus require a high electric current to effectively shim the B0 field. Furthermore, large built-in B0 shimming coils are difficult to dynamically control due to their size, and are not interchangeable from patient to patient.
A second conventional approach to addressing the inhomogeneity of the static B0 field is to use a small dedicated shimming coil placed just outside an RF coil used to image a particular anatomy. This second approach requires less electrical current than the first approach, but, since most commercial MRI systems already have the built-in B0 shimming coils, this reduction in electrical current may be offset by that used by the built-in B0 shimming coils. Additionally, this approach requires a large number of shimming coils that will often completely cover the target anatomy, resulting in a claustrophobic situation for the patient being imaged. Thus, this approach is less than optimal in a clinical situation.
A third conventional approach is to use each RF coil element in an RF coil (e.g., a transmit/receive (Tx/Rx) coil) as a B0 shimming element. In this approach, a shimming current is allowed to flow in the same current path of the Tx/Rx coil, and the shimming current of each RF coil element is controlled by the MRI system to achieve an improved B0 field homogeneity. This third approach improves on the second approach by eliminating some of the excess structure required in the second approach, thereby reducing the cost. Additionally, since the B0 shimming coil shares the same current path as the RF coil, the opening of the RF coil can be kept unchanged, thus mitigating the claustrophobic conditions created by the second approach. However, this approach has drawbacks. First, this approach requires its own dedicated Tx power amplifiers for use in a parallel Tx (pTx) application. Alternately, this approach needs a higher RF power for the associated whole body coil (WBC) used by the MRI system in a non-pTx application.
An MRI system may include two kinds of MRI RF coils. The first kind of MRI RF coil is a transmit (Tx) coil. A Tx coil, while operating in Tx mode, transmits high power RF energy into the anatomy of the subject being imaged to excite nuclei spins in the tissue being imaged. The second kind of MRI RF coil is a receive (Rx) coil. An Rx coil, while operating in Rx mode, detects weak signals from nuclei spins of the anatomy being imaged. A conventional MRI system uses a built-in WBC as a Tx coil. In a conventional MRI system, due to the geometric size of the WBC, the WBC applies RF energy to a much larger region of tissue than is required to image a given region of interest. For example, when a head scan is performed and a WBC is used, not only the head, but also the shoulders and chest also receive a high level of RF energy. This creates a high level specific absorption rate (SAR) issue which limits the clinical utility of MRI systems that use a conventional WBC/Rx coil approach. As a result, a local Tx coil is frequently used to mitigate the SAR problem.
A local Tx coil is designed to apply RF energy into only the anatomy being imaged. There are two conventional approaches to transmitting energy from a power source to a local Tx coil. A first conventional approach is to use a direct connection between the power source to the Tx coil using wires. A direct connection using wires is energy efficient because the energy loss in the connection wires is trivial. A disadvantage of direct connection using wires is that dedicated wiring is required, which increases the cost and complexity of the coil.
A second conventional approach to transmitting energy from a power source to a local Tx coil is to use inductive coupling. In the inductive coupling approach, a primary coil is directly connected to a power source. The primary coil may be a WBC or another large coil. The primary coil is a resonant LC circuit. A smaller second coil (i.e., a local coil) is also used. The second coil is another resonant circuit and is inductively coupled to the primary coil. Thus, energy can be transferred from the primary coil to the second coil. The second coil can be used to excite nearby anatomy more efficiently than the WBC because the second coil is smaller and closer to the nearby anatomy than the WBC. Compared to the first approach using a direct connection with wires, the inductive coupling approach may be less energy efficient than direct wiring but is still more efficient than a conventional WBC. One benefit of the inductive coupling approach is that no special wiring is required. However, conventional inductive coupling approaches require the use of multiple coils. For example, a conventional inductively coupled knee coil uses two layers of RF coils. The first (inner) layer includes a plurality of Rx coil elements which detect signals from the anatomy while operating in Rx mode, and which are decoupled from the transmitting field while operating in Tx mode. The second (outer) layer is typically a standard birdcage coil that inductively couples to a WBC to create a local amplified transmitting field in Tx mode and which is disabled in Rx mode. However, this conventional inductively coupled dual layer coil has drawbacks. For example, all the individual Rx coil elements in a conventional dual layer coil need associated circuits for decoupling the Rx coil and the local Tx coil while operating in Tx mode. Conventional inductively coupled dual layer coils also require circuits for switching off the Tx coil while operating in Rx mode, which requires complex and expensive control circuitry. This leads to complex and expensive coils. These multiple decoupling circuits and complex control circuits can also decrease the signal to noise ratio (SNR), thereby reducing image quality. Furthermore, the outer layer, by its proximity to the inner layer, will create additional noise when the inner layer is operating in Rx mode.
A challenge in implementation of single-layer MRI RF coils is the absolute control of the SAR in any given scenario. Since a single-layer MRI RF coil is inductively coupled to the WBC, SAR patterns may be unpredictable. This unpredictability may increase if a single-layer MRI RF coil array element that is part of a single-layer MRI RF coil array is damaged.